Biological detector and method

ABSTRACT

A biological detector includes a conduit for receiving a fluid containing one or more magnetic nanoparticle-labeled, biological objects to be detected and one or more permanent magnets or electromagnet for establishing a low magnetic field in which the conduit is disposed. A microcoil is disposed proximate the conduit for energization at a frequency that permits detection by NMR spectroscopy of whether the one or more magnetically-labeled biological objects is/are present in the fluid.

This application is a continuation of U.S. patent application Ser. No.13/776,580, filed Feb. 25, 2013, issued as U.S. Pat. No. 8,698,494,which is a continuation of U.S. patent application Ser. No. 12/720,499,filed Mar. 9, 2010, issued as U.S. Pat. No. 8,384,381, which is acontinuation of U.S. patent application Ser. No. 11/894,597 filed Aug.21, 2007, issued as U.S. Pat. No. 8,339,135, which claims the benefit ofU.S. Provisional Application No. 60/839,006, filed Aug. 21, 2006, all ofwhich are incorporated herein by reference in their entirety.

GOVERNMENT INTEREST STATEMENT

This invention was developed under Contract DE-AC04-94AL85000 betweenSandia Corporation and the U.S. Department of Energy. The U.S.Government has certain rights in this invention.

FIELD OF THE INVENTION

The invention relates to a NMR-based biological detector and detectionmethod that involves NMR detection of magnetic nanoparticle-labeledbiological objects using a microcoil.

BACKGROUND OF THE INVENTION

Nuclear magnetic resonance (NMR) spectroscopy is widely used for thereal-time identification of chemical compounds in solids, liquids, andgases because it can easily detect and characterize all components ofmixtures without requiring separations. Unfortunately, standardhigh-resolution NMR spectroscopy is not useful for directly detectingdilute biological objects, such as tumor cells, bacteria, bacterialtoxins, or viruses, in fluid samples. The weak signals from the analytesin the dilute species are lost against the much stronger backgroundwater signal. Even if the dynamic range challenge is met by suppressingthe bulk water signal or concentrating the dilute species, the rapidtransverse relaxation characteristics of macromolecular, viral, orcellular samples renders their direct detection by NMR difficult.

Recent developments involving superparamagnetic iron oxide nanoparticles(SPIONs) have, however, supplied the basis for new applications of NMRwith high sensitivity and specificity for the detection and quantitationof dilute biological materials in fluids, such as cancer cells in bloodor urine samples, or bacterial contaminants in food products or drinkingwater.

SPIONs are enjoying significant uses as biological contrast agents forNMR imaging in human clinical medicine. Furthermore, these nanoparticlescan be coupled with biologically specific recognition ligands to targetepitopes involved in diseases, like cancer. The her-2 protein, forexample, is over-produced in many breast cancers and has been thesubject of successful NMR imaging experiments where cells displayingthis protein have been specifically imaged by means of SPIONs labeledwith anti-her-2 antibodies. The image contrast effects due to SPIONs,which are typically embedded in larger beads, rely on the enhancement ofthe relaxation rates of water molecules surrounding the beads. Themagnetic field gradient from a single, micron-sized magnetic bead hasbeen shown to influence the relaxation time T*₂ of the surrounding waterwithin a voxel approximately 100 μm on a side (a volume of 1 nL), whichis about 1000 times larger than that of a single cell. Thus, for a smallbiological object bound to a magnetic bead in water, the change in theNMR signal caused by the presence of the object is greatly amplified bythe effect of the magnetic bead on the surrounding water.

In recent years, significant advances in the development and fabricationof microcoils (size<1 mm) for NMR have continued. Both planar surfacemicrocoils and solenoidal microcoils have been developed. To enhancesensitivity for tiny samples, much of the work with microcoils hasutilized the high fields produced by strong superconducting magnets.

SUMMARY OF THE INVENTION

The invention provides in an illustrative embodiment an NMR-basedbiological detector and detection method that involve detection of oneor more magnetic nanoparticle-labeled biological objects in a fluid,such as water, contained in a fluid-receiving conduit using a microcoiland a magnetic field generator, such as one or more low field permanentmagnets or electromagnets, to establish a relatively low magnetic fieldwith energization of the microcoil at a frequency that permits detectionby NMR of one or more biological objects present in the fluid.

In an illustrative embodiment of the present invention, the microcoilhas an inner diameter of about 50 to about 550 microns, preferably about75 to about 125 microns, and even more preferably about 100 microns. Themicrocoil can comprise a solenoid-shaped or a flat, planar shapedmicrocoil. The solenoid-shaped microcoil preferably comprises a metallicwire microcoil wound on a tubular microconduit to reduce cost of thedetector. The fluid in the tubular conduit is disposed in a magneticfield of about 0.5 to about 1.5 T established by one or more permanentmagnets or electromagnets. The microcoil on the conduit can be mountedon a ceramic chip substrate to provide a compact assembly. Amicrofluidic chip also can be used to this end.

The NMR-based detector can provide capability of performing routinerelaxation time measurements and low-field spectroscopy for thedetection of dilute concentrations of magnetic nanoparticle-labeledbiological objects in fluids. Such biological objects include, but arenot limited to, cancer cells in blood or urine samples, bacterialcontaminants in food products or drinking water, and biological warfareagents in aqueous media.

The present invention is advantageous in a preferred embodiment inproviding a microcoil together with one or more compact permanentmagnets with benefits of reduced cost, maintenance, and spacerequirements of the NMR-based detector as well as portability thereof.

Other advantages of the invention will become apparent from thefollowing detailed description taken with the following drawings.

DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic view of an NMR-based detector in accordance withan illustrative embodiment of the invention.

FIG. 2 is a flow chart showing a particular illustrative sequence ofsteps for fabrication of a microcoil and connection to electrical leads.

FIG. 3 is a perspective view of a microcoil made pursuant to FIG. 2 andused in testing described below in the Example. A 0.5 mm scale bar alsois shown.

FIG. 4 is a plan view of the microcoil on the fluid sample conduitmounted on a ceramic chip substrate with electrical connectionspartially shown. A 4.0 mm scale bar also is shown.

FIGS. 5A and 5B illustrate different probe circuits for tuning thelow-inductance microcoil at low frequency.

FIG. 6 shows determination of the Π-pulse width in the microcoil from awater sample. The spectral intensity is plotted as a function ofexcitation pulse width. Transmitter power was 0.25 mW. The line is asine wave fit to the data wherein the Π-pulse length given by this fitis 397±4 μs (microsecond).

FIG. 7 shows an absorption spectrum of a sample of de-ionized water,calculated by Fourier-transforming the FID (free-induction decay) from asingle Π/2 pulse of width 200 μs. The full width at half maximum is 2.5Hz, and the line is nearly Lorentzian, as shown by the left inset. Theright inset shows the FID. The time domain data were acquired at 200 μsper point and then digitally filtered to 6.4 ms per point (decimated by32). The signal to noise ratio, measured as the initial FID amplitudedivided by the standard deviation of the baseline noise, is 137. Thesmall peaks near −1.7 and +1.1 ppm (marked by *) are 60 Hz sidebands ofthe main peak; they result from gain variations in the receiver used.

FIG. 8 is a spectrum of 100% ethanol taken with 64 FIDs, 8192 pts, 100μs per point, 10 kHz filters (the lowest available), decimated by 4,with a 5 s (second) relaxation delay. The parameters for a fit of thespectrum to the sum of three Gaussians are listed in Table 1.

FIGS. 9A, 9B, and 9C show measurement of the longitudinal relaxationtimes for three samples of water in the microcoil. FIG. 9A is data fromwater doped with GdDTPA with T₁ determined being 65±4 ms. FIG. 9B isdata from de-ionized water with T₁ determined being 1.0±0.2 s. FIG. 9Cis data from a 1:10 dilution of Dynabeads in de-ionized water with therelaxation time T₁ determined being 0.64±0.17 s.

FIG. 10 shows effect of SPIONs (Dynabeads) on the transverse (T₂*)relaxation time of water in the microcoil. The first 300 ms of the freeinduction decays for the de-ionized water sample, as well as watersamples containing 10, 100, and 1000 beads/nL (not labeled), are shown.For the first three samples, a single scan, digitally filtered to aneffective acquisition rate of 400 μs per point, is shown. For the 1000beads/nL sample, 16 scans were averaged together. The inset shows the ¹HNMR spectra from the FIDs, showing both the increase in line width andthe shift to lower frequency due to the presence of the magnetic beads.

FIG. 11A shows change in 1/T₂* due to the presence of 1 μm magneticbeads as a function of concentration. The solid squares are dataobtained using a 264 nL microcoil, while the open circles are dataobtained from the same bead solutions in 5 mm NMR tubes using aconventional probe. The straight line, drawn as a guide to the eye, hasa slope of roughly ⅔ on this log-log plot, indicating that ΔR₂* isproportional to the ⅔ power of concentration over this range. FIG. 11Bshows relaxivity r₂* of the magnetic beads as a function ofconcentration. Per bead, the enhancement of 1/T₂* increases withdecreasing concentration.

DETAILED DESCRIPTION OF THE INVENTION

The invention provides an NMR-based biological detector and detectionmethod for detection of one or more magnetic nanoparticle-labeledbiological objects in a fluid, such as water. The fluid can be containedin a closed or open microconduit, such as a capillary tube or open-sidedmicrochannel on or in a substrate (microfluidic chip), or other fluidsample holder. The fluid can be introduced to the conduit by injectionusing a fluid sample syringe for example, by capillary action using acapillary tube as the conduit, under pressure by a micropump forexample, and/or any other technique and can be static in the conduit orcan flow through the conduit during practice of the invention. TheNMR-based detector includes a microcoil and a magnetic field generator,such as one or more low field permanent magnets or electromagnets, toestablish a magnetic field of about 0.5 to about 1.5 T (Tesla) withenergization of the microcoil at a frequency that permits detection byNMR of one or more labeled biological objects present in the fluid inthe conduit. The NMR-based detector can provide capability of performingroutine relaxation time measurements and low-field spectroscopy for thedetection of dilute concentrations of magnetic nanoparticle-labeledbiological objects in biological fluids such as blood, saliva, and serumas well as aqueous fluids such as drinking water and aqueous industrialwaste or effluent streams or spills. Such biological objects include,but are not limited to, cancer cells in blood or urine samples,bacterial contaminants in food products or drinking water, andbiological warfare agents in aqueous media.

Referring to FIG. 1, an NMR-based detector pursuant to an illustrativeembodiment of the present invention is schematically shown and includesa fluid-receiving conduit 10 in the form a capillary tube 10 a thatresides in a gap G between opposite polarity ends N, S of first andsecond permanent magnets 12 that provide a magnetic field of about 0.5to about 1.5 T in the gap. The permanent magnets 12 are disposed onrespective soft iron, steel, or other ferrous support members 14 byadhesive or any fastening technique. Soft iron flux-carrying members 16are shown disposed between the support members 14 to complete themagnetic flux path therebetween of the magnet assembly.

The permanent magnets 12 can comprise cylindrical shaped commerciallyavailable SmCo, NdFeB or other low field permanent magnets that providea magnetic field in the range of about 0.5 to about 1.5 T. For example,suitable SmCo and NdFeB permanent magnets are available from Neomax,Osaka, Japan. For purposes of illustration and not limitation, suchpermanent magnets can have a diameter of 2 inches and length of 2 inchesand provide a gap of 0.05 inch width in which the conduit 10 isdisposed. The permanent magnets can have any shape in practice of theinvention. One or more permanent magnets, or a single C-shaped orsimilar shaped permanent magnet having integral opposite polarity endsat a gap can be used in practice of the invention.

Use of one or more low field compact permanent magnets with themicrocoil 20 to be described below provides benefits of reduced cost,maintenance, and space requirements for the NMR-based detector as wellas imparts portability to the detector.

However, the present invention also can be practiced using one or moreelectromagnets to provide a low magnetic field of about 0.5 to 1.5 T inlieu of the one or more permanent magnets described above. Anelectromagnet that can be used can comprise a high quality solenoid, oran iron-core magnet with polished pole faces.

Referring to FIG. 1, a microcoil 20 is shown disposed about the outerperiphery of the conduit 10. In particular, the microcoil 20 is showndisposed about and on the periphery of the capillary tube 10 a in thegap G. The microcoil 10 is connected by electrical leads 11 to a tuningcircuit 30 described below in the Example to energize the microcoil at afrequency that permits detection by NMR of one or more magneticnanoparticle-labeled biological objects present in the fluid. Themicrocoil 20 can be energized to provide the capability of performingroutine relaxation time measurements and low-field spectroscopy for thedetection of dilute concentrations of one or more magneticnanoparticle-labeled biological objects in the fluid in the conduit 10.

The solenoid-shaped microcoil 20 illustrated in FIG. 1 can have an innerdiameter of about 50 to about 550 microns, preferably about 75 to about125 microns, and even more preferably about 100 microns determined bythe outer diameter of the capillary tube 10 a when the microcoil isfabricated directly on the tube. For example, the solenoid-shapedmicrocoil 20 illustrated in FIG. 1 can be formed by depositing one ormore metallic layers on the cylindrical capillary and micromachining thelayers to form a solenoid-shaped microcoil as described in the Examplebelow. Alternately, the solenoid-shaped microcoil 20 can be formed bywinding a metallic wire (e.g. a 50 gauge copper wire) directly on atubular conduit to reduce cost of the detector. Winding of the metallicwire on the capillary tube is achieved by e.g., using small lathe torotate the tube.

As described below in the Example, the conduit 10, the permanent magnets12 (or one or more electromagnets if used), and the microcoil 20 can bedisposed or mounted on a ceramic chip substrate C (see FIG. 4) toprovide a compact assembly for use in a portable NMR-based detector.

The invention envisions detecting dilute concentrations of one or morespecific magnetic nanoparticle-labeled biological objects in the fluidby performing routine relaxation time measurements and low-fieldspectroscopy. For example, for purposes of illustration and notlimitation, such biological objects include, but are not limited to,cancer cells in blood or urine fluid samples, bacterial contaminants influid food products or drinking water, biological contaminants in wastewater or spills, and biological warfare agents in aqueous fluid media.

The biological objects of the fluid are labeled using specificbiological ligands (e.g. antibodies), which are carried onsuper-paramagnetic or other magnetic nanoparticles detectable by NMR.The nanoparticles can include, but are not limited to,superpara-magnetic iron oxide nanoparticles (SPIONs), or nanoparticlesmade out of cobalt, manganese, nickel, or other small paramagneticmaterials. The surfaces of the nanoparticles typically are partially orfully covered or encapsulated by the specific biological ligand (e.g.antibody) to this end, although other particle surface chemistry may beemployed in practice of the invention to provide desired NMR-detectablebioconjugations with the biological objects to be detected.

The recognition of the biological objects by the magnetic-labeledligands (e.g. antibodies) results in a perturbation of the magneticrelaxation times (T₁, T₂, T₂*) and properties of the fluid (water)molecules in the NMR fluid sample to amplify the recognition event to anextent to permit NMR detection of dilute concentrations of the labeledbiological objects. The magnetic perturbations emanating from thepresence of super-paramagnetic nanoparticles are so strong that only afew, possibly one, biological object may be needed to provide adetectable change in the NMR signal. The recognition of the biologicalobjects by the magnetic-labeled ligands (e.g. antibodies) may or may notresult in nano-self-assembly of the labeled biological objects. That is,the present invention does not require that self-assembly of the labeledbiological objects occur.

The following EXAMPLE is offered to further illustrate the inventionwithout limiting the invention in any way:

EXAMPLE

This Example involves an NMR-based detector having a 550 μm innerdiameter, solenoidal microcoil deposited and micromachined on acapillary tube. Superparamagnetic iron oxide nanoparticles (SPIONs) areshown to measurably change the nuclear magnetic resonance (NMR)relaxation properties of nearby protons in aqueous solution.

Microcoils were fabricated onto quartz tubes each having a length of 2.5cm and a 550 μm outer diameter and 400 μm inner diameter using theprocedure depicted in FIG. 2. The quartz tubes are commerciallyavailable from Vitrocom, Mountain Lakes, N.J. Each tube was cleanedusing hydrogen peroxide followed by acetone and isopropyl alcohol. Afterwet chemical treatment, the tubes were masked on each end (step 1, FIG.2), and the 6.2 mm unmasked center targeted for metal deposition wasetched for 15 minutes using a 100 W 0₂/Ar plasma. The central regionlength was chosen based on the coil design with two 2 mm-long cuffs oneither end. The masked tubes were mounted into conventional individualpin vice fixtures for metal deposition. A stage having eight individualrotation stations contained within a high-vacuum thin film depositionchamber allowed for simultaneous coating of multiple tubes at a constantworking distance of 35 cm. Electron beam evaporation was used to deposita thin Cr layer (200 Å) followed by a relatively thick Au layer (5 μm)around the circumference of the tubes (step 2). Deposition rates werechosen to minimize the stress in the Cr and Au layers. After removal ofthe quartz tubes from the deposition system, the masks were removedusing acetone (step 3), and the tubes were re-mounted into conventionalpin vice fixtures for rotation within the focused ion beam (FIB) system(step 4).

Thirty keV Ga ions emitted from a liquid metal ion source were used toremove (micromachine) the Au/Cr layer in order to define the coil andthe neighboring cuffs 20 c. The ion beam was focused to approximately0.5 μm width using a dual-lens Magnum ion column from FEI Co.,Hillsboro, Oreg., and steered across areas outlined by the operatoruntil all the metal was removed from targeted regions (step 4). Rates ofmetal removal were on the order of 10 μm³/s when using a 20 nA Ga beam.Minimal heat and force accompany FIB bombardment. The secondary electronintensity was monitored during ion bombardment to ensure completeremoval of metal and slight penetration into the quartz.

An example microcoil is shown in FIG. 3 (which includes a 0.5 mm scalebar) with areas removed by the FIB appearing as black lines. Asindicated in FIG. 2, step 4, the sample was rotated by an in-vacuum,single-axis rotary stage and translated by a high precision x-y stagealong the tube axis in order to define a helix. The motion-controlsystem, consisting of an ultra-high vacuum compatible stepper motor(controlled by a Princeton Research Instruments stepper motor unit) anda reduction gear assembly, could orient a sample with 0.25° precision.This FIB method possibly may be used to fabricate microcoils onto muchsmaller tubes having approximately 50 μm outer diameter.

The finished metallic coil used in this Example (FIG. 3-4) had 28 turnsover a length of 2.1 mm. The coil conductors were 65 μm wide with a gapbetween turns of 10 μm. The sample detection volume within the NMRmicrocoil was 264 nL. The filling factor was (400/550)²=53%. On the 2-mmlong metallic cuffs of the coil, the FIB removed a 10-μm wide lineparallel to the tube axis in order to interrupt conduction. Thesecondary electron detector within the FIB system also enabledregistration of the coil turns. The direct current resistance (measuredusing a Fluke model 179 resistance measurement unit) of themicromachined coil was found to be 5.42 Ohms. The resistivity of theevaporated Au is 2.898 pOhm-cm (measured on a flat substrate), somewhathigher than bulk Au. Using this value and the geometry of the microcoil,a DC resistance of 4.3 Ohms was calculated. This differs from themeasured resistance, perhaps due to contact resistance in the silverepoxy used to attach the coil to the circuit board. The microcoilinductance was calculated to be 93 nH.

The microcoil/capillary tube were packaged on a ceramic chip substrate Ccomprising DuPont™ Green Tape™ Low Temperature Co-Fired Ceramic (LTCC)material available from DuPont Microcircuit Materials, Research TrianglePark, N.C. The chip substrate had dimensions of 20 mm by 60 mm. The chipsubstrate had been previously plated with alloyed gold (Au—Pt)co-firable material (DuPont 5739) solder leads 11 a as shown in FIG. 4which are connected to leads 11. The microcoil was secured to thealloyed gold leads by means of silver-containing epoxy (FIG. 4 andschematically shown as step 5, FIG. 2), the alloyed gold leadssupporting the microcoil on the chip substrate above an opening OP (FIG.4) in the substrate. The opening assured that the microcoil did notcontact the supporting platform and prevented distortion or damage tothe very thin metal layer. Mounting of the microcoil on the ceramic chipsubstrate in this way allowed manipulation of the microcoil andattachment of the capillary tube on the substrate without damaging thecoil.

′H NMR measurements, at a resonant frequency of 44.2 MHz, were performedusing a MRTechnology console (Tsukuba City, 300-2642 Japan), and a 1.04T (Tesla) NEOMAX permanent magnet assembly comprising a NdFeB permanentmagnet assembly providing a 2 inch gap. The microcoil on the chipsubstrate was inserted in the gap between the magnet pole faces with thechip substrate supported in the gap by plastic spacers for the NMRmeasurement. A smaller 1 Tesla permanent magnet suitable for use in aportable microcoil NMR device can be fabricated.

The transmitter pulses were output directly from the console, without aconventional radiofrequency power amplifier, because only 0.25 mW ofpower was required to produce a B₁ field of 0.3 G (vide infra). Ethanol(100%) was purchased from AAPER (Shelbyville, Ky.). Spin-lattice ¹H T₁values were obtained, using a standard inversion-recovery sequence, froma Gd-DTPA-containing water sample (Gd-DTPA isgadolinium-diethylene-triamine-pentaacetate), from a sample of magneticbeads in water, and from a sample of de-ionized water.

Magnetic beads (Dynabeads; MyOne Streptavidin) were purchased from DynalInc. Each magnetic bead consists of thousands of 8-nm diametersuperparamagnetic iron oxide particles, uniformly dispersed in apolystyrene matrix, and coated with a thin layer of polymer and amonolayer of streptavidin which served as a bonding agent onto whichbiotinylated antibodies could be attached. The beads are 26% Fe byweight (about 10% Fe by volume) with an average diameter of 1.05±0.10μm. The stock solution has a stated bead concentration of between 7×10³and 1.2×10⁴ beads per nL (equivalent to about 2.6 mg Fe/ml). NMR sampleswere prepared by diluting the same batch of stock solution withde-ionized water by factors of 10, 100, and 1000 to produce nominalconcentrations of 1000, 100, and 10 beads per nL introduced to thedetector tube by supply syringe. The relaxation time T*₂ was determinedby collecting a single free-induction decay (FID) and fitting theresulting spectrum with a Lorentzian, unless noted otherwise. Therelative shift of the NMR frequency of water caused by themagnetic-labeled beads was determined by measuring the resonancefrequency of each solution in a 5 mm NMR tube in a conventional coilrelative to a separate tube of deionized water. To avoid errors due tofield drift of the permanent magnet, each frequency shift measurementwas performed by switching several times between the bead solution and adeionized water sample during a period when the frequency drift wasconfirmed to be less than 1 Hz/min.

The 93 nH inductance of the 550-μm outer diameter microcoil describedabove could reach resonance at 44.2 MHz with a variable capacitor ofreasonable size. However, since use of much smaller microcoils at coilouter diameters nearer the 50 μm outer diameter are envisioned,described below are tuning circuits for tuning such smaller microcoilsto resonance at the 44.2 MHz resonant frequency or less to detect waterwith a spectral resolution of 2.5 Hz in a 1.04 Tesla permanent magnet.

Such a tuning circuit involves an auxiliary tank circuit withconventional scale capacitors and to connect the microcoil to it. Thekey parameter of the microcoil described above that guided the design ofthis tuning circuit was its very high coil resistance. Optimization of atuned circuit's SNR (where SNR is the signal to noise ratio) is acompromise between maximizing coil efficiency, in terms of the magneticfield produced per unit current in the sample coil, while minimizing theresistive noise. The dominant noise source for the very thin,ribbon-wire shaped microcoils used in this Example was its large coilresistance. Therefore, the introduction of the additional inductor didnot degrade performance, because this extra inductance did notcontribute to the resistive losses.

Two tuning circuits were constructed for use in the Example as shown inFIGS. 5A and 5B. In both cases, the microcoil was mounted by itself in acast aluminum box, while the external tuning inductor L and tuning andmatching capacitors C_(I), C_(M) were mounted in a separate aluminumbox. The capacitors are adjusted to yield an input impedance of 50 Ohmsfor the combined circuit. In the first circuit (FIG. 5A), a quarter-wave(λ/4) (50 Ohm characteristic impedance) cable was used to transform thecoil resistance to a higher value and then placed this transformedimpedance in parallel with the tuning inductor. In this case, the fullresonant voltage was applied to the (transformed) sample coil impedance.In the second circuit (FIG. 5B), the connection between the two parts ofthe circuit was short, and the sample coil and tuning inductor were inseries, so that all of the resonant current flowed through the samplecoil.

The two circuits exhibited nearly identical SNR performance. Allsubsequent measurements were performed with the first circuit (FIG. 5A),because the remote placement of the tuning and matching elements made itmore convenient to work with. The external “tuning” inductor in thiscircuit was 5 turns of 14 gauge bare copper wire, with a calculatedinductance of 0.25 μH, and a calculated resistance at 44.2 MHz of 0.07Ω.Hence, the tuning inductor contributes negligibly to resistive noise;the tuning circuit is therefore as efficient as a conventional circuitmade without the extra tuning inductor. The tuning and matchingcapacitances were both approximately 22 pF. The large value of thematching capacitance resulted from the high losses in the microcoil.Because the Wavetek radio frequency sweeper used in the Example operatesat the milliwatt level, and there was reluctance to subject themicrocoil described above to this power, the Q of the resonant circuitwas estimated by constructing a mockup of the microcoil using robust 36gauge copper wire and a 5 Ω(ohm) resistor. The mockup circuit had a Q ofabout 10, as measured from the halfpower points on the sweeper output.

The nutation performance of the microcoil probe is shown in FIG. 6,where the signal intensity, after an excitation pulse, from a sample ofde-ionized water, is plotted as a function of pulse width α. The datafollowed a typical sin(α) curve, indicating uniform sample excitation bya homogeneous RF field. The Π-pulse width, determined from fitting thesine curve, was 397±4 μs. The transmitter amplitude was 0.32 V(peak-to-peak), corresponding to a power into 50Ω of only 0.25 mW. AΠ/2-pulse time of 200 μs corresponds to an RF field strength of 0.3 G,which is produced in the coil by a current of 1.8 mA.

The free-induction decay (FID) and spectrum of deionized water in themicrocoil are shown in FIG. 7. The spectrum has a full-width at halfmaximum (FWHM) of 2.5 Hz (0.056 ppm) and is reasonably well-fit by aLorentzian, as shown in the left inset. (At 55% and 11% of maximum, thewidths are 2.3 Hz and 8.7 Hz, respectively.) The SNR after a single Π/2pulse was found to be 137 (ratio of FID amplitude to rms baselinenoise). The small sidebands at ±60 Hz were presumably due to gainmodulations in the receiver amplifiers used, caused by 60 Hz ripple.(Sidebands ±120 Hz were also observed.) FIG. 8 shows the NMR spectrum ofa sample of 100% ethanol, calculated from 64 FIDs acquired with a 5 srepetition time. Peaks are seen at δ=1.2, 3.7, and 5.5 ppm,corresponding to the CH₃—, CH₂—, and —OH protons, respectively, with thecorrect relative amplitudes of 3:2:1 (Table 1). Note also theobservation of the approximate 7 Hz J-coupling for the methyl group, andthe smaller couplings for the methylene and hydroxyl protons, indicatingthat the frequency drift over the 5-minute experiment was less than 3Hz. For both the water and the ethanol experiments, only the X, Y, and Zgradients were shimmed because higher order shims were not available.

TABLE 1 Fit of the ethanol spectrum to the sum of three Gaussians δ(ppm)Multiplicity Amplittude 1.2 [1.2]°  [3] 3.0 [3] 3.70 [3.65]  [4] 1.9 [2]5.48 [5.275] [1] 1.1 [1] ⁰ The standard values are shown in squarebrackets

To test the ability of the microcoil to measure spin-lattice relaxationtimes, three different water samples were used; the first sample wasdoped with Gd-DTPA to shorten the T₁ to around 70 ms, the second sampleconsisted of pure de-ionized water, and the third sample containedmagnetic beads (at a concentration of 1000 beads/nL) in de-ionizedwater. In all cases, a single scan was acquired at each recovery time.The results (FIG. 9A, 9B, 9C) show that relaxation times can beaccurately measured for both shorter (65 ms) and longer (0.6 and 1.0 s)T₁ values with a standard inversion-recovery pulse sequence. The 397 μsΠ-pulse gave clean inversion of the magnetization for all samples.

In FIG. 10, the signal detected from deionized water and three differentdilutions of the stock Dynabead solution, corresponding to 1000, 100,and 10 beads/nL are compared. The magnitude of each FID is shown, sothat they all appear as if they were on resonance. The data are acquiredafter a single Π/2 pulse, digitizing at 100 μs per point (200 μs perpoint for the deionized water). The data were digitally filtered toachieve an effective digitization time of 400 μs per point. For the 1000beads/nL sample, 16 FIDs were averaged together; the other data are eacha single FID. The beads have two effects on the water spectral peak: thepeak broadens and shifts to lower frequencies as the concentration ofbeads increases. The reduction in T*₂ is apparent in the FIDs. The insetof FIG. 10 compares the spectra of the four solutions and shows both thelinebroadening and the shift to lower frequency caused by the beads.Data for a 1 bead/nL sample (not shown) were indistinguishable from thedeionized water data.

The solid symbols in FIG. 11A give the observed change in 1/T*₂ (ΔR*₂)due to the presence of the beads, as a function of bead concentration,C. Here, ΔR*₂=R*_(2 bead solution)−R*_(2 water) and R*₂=ΠΔf, where Δf isthe FWHM in Hz of the Lorentzian line fit to each spectrum in FIG. 10.Note that both axes in FIG. 11A are logarithmic; the straight line(drawn as a guide to the eye) has a slope of roughly ⅔ indicating thatΔR*₂ is proportional to C^(2/3) over this range of bead concentrations.The relaxivity r*₂(=ΔR*₂/C) is therefore not a constant, but decreaseswith increasing concentration as shown in FIG. 11B.

Because magnetic field gradients can cause motion of the magnetic beadswith respect to the fluid, it was not clear a priori that theconcentration of beads delivered to the microcoil would be the same asthe concentration in the supply syringe. Indeed, the measured T*₂ ofbead solutions in the microcoil was observed to decrease over time ifthe bead solution was allowed to sit motionless in the coil over severalminutes, suggesting that the spatial distribution of the beads waschanging, due to clustering, settling, or migration out of the coil.Thus, in order to validate the microcoil results, the T*₂ of the samebead solutions (1000, 100, and 10 beads/nL) and deionized water incapped 5 mm NMR tubes using a conventional probe in the same magnet.Each measurement was performed within 20-30 s after shaking the tube tohomogenize the bead solution, and the tube was immediately extractedafterwards to visually confirm that the beads had not settled during themeasurement. (Shimming was performed on the deionized water, and asample holder was used to position the other 5 mm tubes identically, toavoid the need to re-shim. Repeatedly placing the same sample in theprobe using this holder gave linewidths that were reproducible to ±5Hz.) Migration of the beads was similarly observed in the 5 mm tubes(both visually and as an increase in T*₂ over time) if the samples wereallowed to sit in the magnet for longer time periods. The ΔR*₂ valuesmeasured for the bead solutions in 5 mm tubes (open symbols in FIG. 11A)are in good agreement with those obtained for the same concentrations inthe microcoil, indicating that the expected concentrations weredelivered to the microcoil.

Thus, in summary, the nutation performance of the microcoil wassufficiently good so that the effects of magnetic beads on therelaxation characteristics of the surrounding water could be accuratelymeasured. The solution of magnetic beads (Dynabeads MyOne Streptavidin)in deionized water at a concentration of 1000 beads per nL lowered theT₁ from 1.0 to 0.64 s and the T*₂ from 110 to 0.91 ms. Lowerconcentrations (100 and 10 beads/nL) also resulted in measurablereductions in T*₂, indicating that low-field, microcoil NMR detectionusing permanent magnets can be used as a high-sensitivity,miniaturizable detection mechanism for very low concentrations ofmagnetic beads in biological fluids.

The tuning circuit described above is capable of tuning an arbitrarilysmall inductance at a frequency compatible with a permanent magnet,coupled with the 550 μm microcoil, allows spectroscopic and relaxationmeasurements using less than 1 mW of radiofrequency power. (This lowpower requirement further aids in making the NMR detector portable.) Theline widths for deionized water are adequate for the detection ofmagnetic beads in water at a concentration of 10 beads/nL. The coil usedfor these proof-of-principle measurements is not optimized in size forNMR sensitivity, as discussed further below. However, the above resultsindicate that this approach will allow the detection of very dilutebiological species, perhaps as rare as a single cell or molecule labeledwith a single magnetic bead.

The challenge of achieving this detection sensitivity can be discussedquantitatively in light of the data of FIGS. 10 and 11. In a portabledetector system pursuant to the invention, a fluid containing verydilute, magnetically labeled biological objects flows through aapproximate 1 nL volume coil while the FID is monitored. The challengeis to detect the difference between the FID of the background fluid andthe same fluid containing one magnetic bead (nanoparticle) within thecoil volume. Considering FIG. 10, the change in T*₂ of water can bereadily detected due to 10 beads/nL, or roughly 3000 magnetic beads(nanoparticles) in the Example microcoil (264 nL volume). Achievement ofa similar T*₂ for deionized water (approximate 100 ms) and adequate SNRin a microcoil with a 1 nL sample volume permits detection of 10 beads(nanoparticles).

Extrapolating the straight line in FIG. 11A indicates that the ΔR*₂ ofone bead in a 1 nL volume is about 8 s⁻¹, which would have caused anincrease in the linewidth of water in the Example microcoil from about 3to about 6 Hz. This increase should have been detectable given the highSNR. That such a change in line width due to the 1 bead/nL solution wasnot detected suggests that the ΔR*₂ for this concentration is lower thanthat predicted by extrapolating the straight line in FIG. 11A. Atheoretical treatment of dipolar broadening of the NMR line due todilute magnetic impurities indicates that the linewidth will beproportional to C^(1/2) at higher concentrations and will be linear in Cat lower concentrations. The above slope of ⅔ suggests that the exampleconditions are in the transition region between these two limits suchthat a higher slope at lower concentration should be expected, resultingin a predicted value of ΔR*₂ lower than 8 s⁻¹ at 1 bead/nL. Hence, thedetection of a single 1-μm Dynabead in a 1 nL coil will requireachievement of an even narrower line width, while at the same timedetecting adequate signal strength.

A 100 μm diameter coil (1 nL) will give substantially less signal thanthe Example 264 nL microcoil due to the reduced sample size. Thus onemust consider whether such a coil will have sufficient SNR to detect 10beads in its 1 nL volume. In the “large” microcoil data in FIG. 10,detection sensitivity can be maximized by integrating the FIDs, say from50 to 300 ms, which is roughly equivalent to applying strong digitalfiltering. These integral values are 397 and 122 (arbitrary units), forthe water and 10 beads/nL data, respectively. The uncertainty in thesevalues is 3, which corresponds to a signal to noise ratio (SNR) of 133for determining the amplitude of the water signal. The smaller 1 nLvolume coil will have much less signal, but also less noise (due to itslower resistance). For microcoils in the limit where skin depth is smallcompared to wire size (which is not quite true for our coil), the SNRper unit volume scales as the inverse of the coil diameter. Hence, theabsolute SNR scales as the square of the linear dimension of the sample.The invention envisions scaling the sample and coil dimension down byroughly a factor of 6, so that the SNR in the determination of theintegrated water signal amplitude will be about 3.7. Hence, the 1 nLcoil will require that the beads (nanoparticles) change the area underthe FID of the background water by at least 25%. In this Example, aconcentration of 10 bead/nL caused a 70% change in the integrated signalfrom 50 to 300 ms, and therefore may remain detectable in the 1 nL coil,provided that a similar background water T*₂ is achieved.

While the Example microcoil described above is already capable ofdetecting the presence of as few as 3000 magnetic beads (nanoparticles),it can be further optimized for maximal SNR performance for operation at44.2 MHz. The thickness of the coil “wire” is much less than a skindepth, which raises the resistance of the coil without providing anyimprovements in signal detection. The width of the “wire” is much morethan a skin depth, so that it may be possible to increase the number ofturns per unit length and gain in coil sensitivity without suffering anullifying increase in resistance. Careful attention to the geometricaldesign of the smaller microcoil, should improve the SNR above theestimate of about 3.7 based on this Example. SNR performance will beenhanced by reducing the coil resistance, which is higher than expectedin Example ion-milled microcoil. Improving the line width of thebackground fluid places a lower demand on the SNR performance. The useof susceptibility matching (either in the choice of evaporated metals orvia a matching fluid) and the reduction of the filling factor (byincreasing the relative wall thickness in the capillary tube) mayimprove the line widths in smaller coils. In addition, the permanentmagnet used in the Example is not very homogeneous and only first ordershims are available; a more homogeneous applied field may be required toachieve narrower lines. Optimization of the coil can also includecomparisons of both the SNR and line width performance of ion-milledcoils to other types of microcoils, such as copper wire-wound coils.Some compromise between line width and sensitivity may provide the bestopportunity for detecting single biological objects.

The surface of a single cancer cell (about 10 μm in diameter) can bearupwards of 10⁵ binding sites (antigens) for a particular antibody andcan accommodate up to 400 one micron diameter magnetic beads(nanoparticles), assuming monolayer coverage and random close packing.Thus, sensitivity to 10 beads would already be adequate to detect singlemagnetically labeled cells. On the other hand, bacterial toxin molecules(e.g., botulism toxin) are much smaller and would accommodate only oneor a few beads, requiring single-bead detection sensitivity. Hence,single-bead sensitivity is envisioned by practice of the presentinvention.

So far in the Example, the detection limits have been based onmeasurements of a particular type (Dynabeads) and size (1 μm) ofmagnetic bead (nanoparticle). Larger magnetic beads (having largermagnetic moments) are available and will permit an increase in therelaxivity of a single bead and further lower the detection limit.Assuming that a background T*₂ was at least 100 ms, the ΔR*₂ for asingle 1.63 μm bead in a 1 nL volume to be at least 60 s⁻¹, which shouldbe readily detected using a 1 nL microcoil with a background water T¹ ₂of 100 ms and a SNR of about 3. Even larger beads (e.g., 2.8-μm and4.8-μm Dynabeads) are commercially available, and may be used, ifnecessary, to further enhance the ability to detect a single magneticbead in an NMR microcoil.

Although the invention has been described hereinabove in terms ofspecific embodiments thereof, it is not intended to be limited theretobut rather only to the extent set forth hereafter in the appendedclaims.

What is claimed is:
 1. A method for detecting a target analyte in asample, the method comprising: introducing magnetic particles comprisingtarget-specific moieties to a sample comprising a plurality of targetanalytes under conditions such that at least one cluster of magneticparticles and target analytes is formed; and using a nuclear magneticresonance (NMR) device in order to measure a signal due to a change inrelaxation rate, wherein the relaxation rate changes in the presence ofthe at least one cluster, thereby detecting the target analyte in thesample.
 2. The method according to claim 1, wherein measuring isaccomplished in the presence of other components in the sample.
 3. Themethod of claim 1, wherein the changes are selected from relaxationtimes T1, T2, T2*, or a combination thereof.
 4. The method according toclaim 1, wherein the NMR devices comprises: a magnet configured toproduce a magnetic field; a vessel configured to hold the sample andbeing located within the magnetic field; a radio frequency (RF) coillocated within the magnetic field.
 5. The method according to claim 4,wherein the RF coil wraps around the vessel.
 6. The method according toclaim 4, wherein the magnetic is a permanent magnet.
 7. The methodaccording to claim 1, wherein the target-specific moieties areantibodies.
 8. The method according to claim 1, wherein the sample is abiological sample.
 9. The method according to claim 8, wherein thebiological sample is blood or urine.
 10. The method according to claim1, wherein the magnetic particles are superparamagnetic particles.